1. Field of the Invention
This invention generally relates to a long positron emission tomography detector array, and in particular, to a long positron tomography detector array having different sizes of photomultiplier tubes (PMTs).
2. Discussion of the Background
The use of positron emission tomography (PET) is growing in the field of medical imaging. In PET imaging, a radiopharmaceutical agent is introduced into the object to be imaged via injection, inhalation, or ingestion. After administration of the radiopharmaceutical, the physical and bio-molecular properties of the agent will concentrate at specific locations in the human body. The actual spatial distribution of the agent, the intensity of the region of accumulation of the agent, and the kinetics of the process from administration to eventually elimination are all factors that may have clinical significance. During this process, a positron emitter attached to the radiopharmaceutical agent will emit positrons according to the physical properties of the isotope, such as half-life, branching ratio, etc.
The radionuclide emits positrons, and when an emitted positron collides with an electron, an annihilation event occurs, wherein the positron and electron are destroyed. Most of the time, an annihilation event produces two gamma rays at 511 keV traveling at substantially 180 degrees apart.
By detecting the two gamma rays, and drawing a line between their locations, i.e., the line-of-response (LOR), one can retrieve the likely location of the original disintegration. While this process will only identify a line of possible interaction, by accumulating a large number of those lines, and through a tomographic reconstruction process, the original distribution can be estimated. In addition to the location of the two scintillation events, if accurate timing (within few hundred picoseconds) is available, a time-of-flight (TOF) calculation can add more information regarding the likely position of the event along the line. Limitations in the timing resolution of the scanner will determine the accuracy of the positioning along this line. Limitations in the determination of the location of the original scintillation events will determine the ultimate spatial resolution of the scanner, while the specific characteristics of the isotope (e.g., energy of the positron) will also contribute (via positron range and co-linearity of the two gamma rays) to the determination of the spatial resolution the specific agent.
The collection of a large number of events creates the necessary information for an image of an object to be estimated through tomographic reconstruction. Two detected events occurring at substantially the same time at corresponding detector elements form a line-of-response that can be histogrammed according to their geometric attributes to define projections, or sinograms to be reconstructed. Events can also be added to the image individually.
The fundamental element of the data collection and image reconstruction is therefore the LOR, which is the line traversing the system-patient aperture. Additional information can be obtained regarding the location of the event. First, it is known that, through sampling and reconstruction, the ability of the system to reconstruct or position a point is not space-invariant across the field of view, but is better in the center, slowly degrading toward the periphery. A point-spread-function (PSF) is typically used to characterize this behavior. Tools have been developed to incorporate the PSF into the reconstruction process. Second, the time-of-flight, or time differential between the arrival of the gamma ray on each detector involved in the detection of the pair, can be used to determine where along the LOR the event is more likely to have occurred.
The above described detection process must be repeated for a large number of annihilation events. While each imaging case must be analyzed to determine how many counts (i.e., paired events) are required to support the imaging task, current practice dictates that a typical 100-cm long, FDG (fluoro-deoxyglucose) study will accumulate several hundred million counts. The time required to accumulate this number of counts is determined by the injected dose of the agent and the sensitivity and counting capacity of the scanner.
PET imaging systems use detectors positioned across from one another to detect the gamma rays emitting from the object. Typically a ring of detectors is used in order to detect gamma rays coming from each angle. Thus, a PET scanner is typically substantially cylindrical to be able to capture as much radiation as possible, which should be, by definition, isotropic. The use of partial rings and rotation of the detector to capture missing angles is also possible, but these approaches have severe consequences for the overall sensitivity of the scanner. In a cylindrical geometry, in which all gamma rays included in a plane have a chance to interact with the detector, an increase in the axial dimension has a very beneficial effect on the sensitivity or ability to capture the radiation. Thus, the best design is that of a sphere, in which all gamma rays have the opportunity to be detected. Of course, for application to humans, the spherical design would have to be very large and thus very expensive. Accordingly, a cylindrical geometry, with the axial extent of the detector being a variable, is realistically the starting point of the design of a modern PET scanner.
Once the overall geometry of the PET scanner is known, another challenge is to arrange as much scintillating material as possible in the gamma ray paths to stop and convert as many gamma rays as possible into light. In order to be able to reconstruct the spatio-temporal distribution of the radio-isotope via tomographic reconstruction principles, each detected event will need to be characterized for its energy (i.e., amount of light generated), its location, and its timing. Most modern PET scanners are composed of several thousand individual crystals, which are arranged in modules and are used to identify the position of the scintillation event. Typically crystal elements have a cross section of roughly 4 mm×4 mm. Smaller dimensions and non-square sections are also possible. The length or depth of the crystal will determine how likely the gamma ray will be captured, and typically ranges from 10 to 30 mm. The detector module is the main building block of the scanner.
As discussed above, PET imaging relies on the conversion of gamma rays into light through fast and bright scintillation crystals, and the time pairing of individual events to recreate the location of the annihilation process. These actions require very fast detector and electronic components, but no detector design and electronics can compensate for the amount of light lost in the path from the scintillation event to the light sensor. The faction of the total amount of light collected over the amount created in the scintillator is a good measure of the efficiency of the design. To maximize the amount of light collected, it is generally better to arrange the light sensor as close as possible to the scintillation crystal and to avoid reflection and other edge effects. This forces the arrangement to be a large array detector with a short distance between the crystal and the sensor.
On the other hand, fast counting requires that multiple events be processed simultaneously, favoring optical isolation between scintillation events, and the creation of smaller detector blocks. However, a PET imaging system includes more than just a counter. In addition to detecting the presence of a scintillation event, the system will need to identify the location of the event. By properly documenting how light is being distributed to the multiple light sensors, it is possible to assign an event location for any given set of sensor responses. In this case though, light needs to be distributed to multiple sensors, contrary to the requirement of the collecting as much light at possible, and to optically isolate all events.
Currently available PET scanners have two main detector module designs. The first type is a large area detector in which an array of crystals that covers the entire axial extent of the cylinder is formed. Several modules are then arranged together to form a cylinder, each module being optically coupled to the next. An array of photosensors (e.g., photomultiplier tubes or PMTs) is placed on the modules and on the interfaces between modules. See the design shown in FIG. 1A, which illustrates a module that includes an array of crystal elements and an array of PMTs. This approach minimizes the number of optical interfaces and boundaries, and ensures excellent light collection. However, this design suffers from larger numbers of sensors being exposed to the light of a single scintillation event, potentially limiting the ability to process events occurring close to each other, as well as limiting the overall counting capacity.
The second design is based upon an optically isolated block having, for example, four PMT sensors, so as to allow for simplified crystal identification. In the design of FIG. 1A, a block element is composed of four photomultiplier tube sensors on an approximately 50 mm×50 mm crystal assembly. In this approach, the crystals extend to the very edge of the array and a relatively thick light guide is therefore necessary to capture enough light from all PMTs to be able to detect the position of the event. A detector is then formed by arranging multiple blocks (e.g., three or four) to fill out the axial extent, and then repeating this pattern to create the overall cylinder. See the designs shown in FIGS. 1B and 1C. The advantages of this approach include greater flexibility (the detector block is potentially fully functional outside of the scanner (meaning that the detector block can be calibrated separately—as opposed to the large area, continuous detector that can only be calibration as part of a complete system—offering advantages for service at the customer site and for manufacturing of the scanner)) and better count capacity due to the potential parallel operation of each module. The disadvantages of this design are the inclusion of a large number of optical surfaces, potentially interfering with efficient light collection, and a more limited set of options for sensor coverage.
In addition to the overall geometry and the design of the detector module, a third major element in defining a modern PET system is the crystal-light sensor assembly. Two main factors influence the design of the light sensor coverage of a crystal array: (1) the sensor layout, and (2) the sensor size. Both factors affect the cost of a scanner of a pre-determined size. In the case of round photomultipliers, only two layout options are possible: rectangular and hexagonal compact arrangements. The hexagonal compact arrangement offers a higher density of coverage, but is essentially incompatible with the small square block design. Existing scanners therefore favor hexagonal compact layout for large-area module designs, and rectangular layout for small-square block designs.
In order to capture as much light as possible from the scintillation process, for the same crystal assembly and optical material, the ratio of photosensor area to the surface of crystal is a good indicator of the detection potential. While the hexagonal arrangement has the best coverage of approximately 90% (π/4·sin(π/6)) in a continuous mode (infinite crystal and PMT arrays), its efficiency to cover square or rectangular areas greatly suffers at the boundaries. A rectangular layout (even with a circular sensor) is certainly more convenient, even with a decrease of coverage efficiency to 78% (π/4).